1. Field of the Invention
The present invention generally relates to a system and method for calibrating and tuning a gamma ray camera (“gamma camera”).
2. Discussion of the Background
Gamma cameras are primarily used by physicians and medical practitioners who specialize in the field of nuclear medicine. Nuclear medicine is a unique medical specialty wherein low-level radioactive materials (called radionuclides or radiopharmaceuticals) are used to generate images of the organs, bones or tissues of the body. The images generated by gamma cameras are primarily used to determine whether organs or tissues are functioning properly.
Typically, the radionuclides or radiopharmaceuticals are introduced orally or intravenously into the body of a patient. Radiopharmaceuticals are specially formulated to collect temporarily in a specific part of the body to be studied, such as the patient's heart or brain. Once the radiopharmaceuticals reach the intended organ, they emit gamma rays that are then detected and measured by the gamma camera.
A typical gamma camera includes a large area scintillation crystal, which functions as a gamma ray detector. The crystal is typically sodium iodide doped with a trace of thallium (NaI(T1)). The crystal converts high-energy photons (e.g., gamma rays and X-rays) into visible light (i.e., lower energy photons). The crystal is positioned to receive a portion of the gamma ray emissions from the radiopharmaceutical in the body.
When a gamma ray strikes and is absorbed in the scintillation crystal, the energy of the gamma ray is converted into flashes of light (i.e., a large number of scintillation photons) that emanate from the point of the gamma ray's absorption in the scintillation crystal. A photo-multiplier tube (PMT), which is optically coupled to the scintillation crystal, detects a fraction of these scintillation photons and produces an output electronic signal (e.g., current or voltage pulse) having an amplitude that is proportional to the number of detected scintillation photons. The gamma ray camera typically has a plurality of photomultiplier tubes placed in a two dimensional array, with the signals from the different photomultiplier tubes being combined to provide an indication of the positions and energies of detected gamma rays.
The scintillation photons emitted from the detector crystal are typically in the visible light region of the electromagnetic spectrum (such as with a mean value of about 3 eV for NaI(T1)). The scintillation photons spread out from the point of emission. A large fraction of the scintillation photons are transported from the point of emission to a light sensitive surface, called the photocathode, of the PMTs. A fraction of the scintillation photons incident on the photocathodes cause an electron to be emitted from the photocathode. The electron, also called a photoelectron, is then electrostatically accelerated into an electron-multiplying structure of the PMT, which causes an electrical signal to be developed at an output of the PMT. The amplitude of the electrical signal is proportional to the number of photoelectrons generated in the PMT during the time period that scintillation photons are being emitted. Thus, after a gamma ray absorption event, the PMT outputs an electrical signal that can be used with other signals from other PMTs to determine the location of the gamma ray absorption event.
The number of scintillation photons producing electrical signals in each PMT is inversely related to the distance of the PMT from the point of gamma ray absorption, or event location. It is because of this relationship that the position of the event can be calculated from the signals of the PMTs surrounding the event location.
Ideally, the signal derived from each PMT should have exactly the same proportional relationship to the distance from the event location as for all other PMTs. In addition to distance from the event location, the amplitudes of the signals derived from each PMT are proportional to two basic parameters: 1) the number of scintillation photons detected by a PMT, and 2) the gain or amplification of the PMT. Therefore, accuracy to which the position of the event location can be calculated depends on these two factors remaining constant over time.
Typically, a gamma camera is tuned prior to its operation so as to ensure that the camera will calculate accurately the positions of event locations anywhere within an area called the field of view (FOV). Common commercial, large FOV gamma cameras have between about 50 and 100 PMTs. A tuning procedure will typically require a number of steps that balance or equalize the signal amplitudes of the PMTs. The gains of the PMTs are adjusted such that the sum of the signals from all the PMTs are approximately equal in response to a fixed energy gamma event, regardless of the location of the event.
To tune a gamma camera according to known protocol, a known pattern of event locations is presented to the camera, usually by placing a mask of precisely spaced holes over the camera crystal, so that event location calculations can be calibrated to provide the known locations fixed by the positions of the holes, where the gamma rays can pass through the mask. The exact tuning and/or calibration steps may be different among cameras produced by different manufacturers. However, once the tuning and calibration steps are complete, the image quality, which depends on the camera's ability to accurately position event locations, depends on the transport of scintillation light to the PMTs and the gains of the PMTs remaining unchanged from the time when the tuning and calibration procedures were performed.
A number of factors can cause a change in either the gain of a PMT or the light collection properties of the camera. PMT gain is a strong function of temperature, counting rate (i.e., the number of event signals per unit time), and the high voltage (HV) power supply regulation. Additionally, PMTs change their gain over time as they age. The light collection from the crystal to the photocathodes of the PMTs can change if the transmissive properties of lightguide surfaces change. For example, the PMTs are optically coupled to a glass or plastic lightpipe using either an optical grease or epoxy. If any of these materials' light transmissive properties change over time, then the transport of scintillation photons to the PMT will change. Additionally, NaI(T1) is a hygroscopic material, and if water vapor reaches the crystal it becomes yellow and the light transmission is diminished.
Different manufacturers have developed and implemented different means to maintain the constancy of PMT gains. These means typically fall into two categories: 1) automatic (i.e., not requiring the user to initiate the process), and 2) user quality control (QC) procedures (i.e., procedures initiated by the user). Generally, a combination of both automatic and QC procedures has been required.
One automatic system, for example, utilizes light-emitting diodes (LEDs) coupled into the photomultiplier tubes to provide a light signal for calibration of each individual tube. A constant fraction of the light emitted by the LED is incident on the light-sensitive photocathode of the PMT. The PMT output signal is checked against a reference that was set at the time of the last calibration. The gain of the PMT is adjusted if the measured signal has strayed from the reference.
This gain calibration technique depends on the light-emitting diodes having a constant light output for each pulse. Light-emitting diodes, however, do not have constant light output as a function of temperature, and may also vary over the lifetime of the diode. Another drawback of this technique of automatic calibration is that the light from the diode is mostly directly incident on the photocathode of the photomultiplier tube. Therefore, the transport of the light through the scintillation crystal, and associated optical elements, is not significantly sampled by the pulse of light from the diode.
User-initiated QC procedures usually require the placement of a radioactive source to uniformly illuminate the camera. The system acquires an appropriate number of events to achieve statistically significant sampling of each event location. A computer program then analyzes the measured energies and/or image of event locations to determine whether or not the system has drifted away from the properly calibrated state. Many variations of this procedure are possible, but all typically require the user to position a source of radioactivity and initiate the computer controlled acquisition and analysis. Additionally, the procedures also typically require the user to remove the collimator from the camera.
Such QC procedures are cumbersome to the user. If they can be initiated at the end of the day, and complete themselves automatically, then the user's time required is minimal. However, radioactive sources that must be left out in a room overnight require institutional procedures for securing the room, logging out the source and returning it in the morning, and prohibiting access to the room by cleaning and unauthorized personnel. Performing QC procedures during working hours reduces available patient imaging time on the system and increases costs because personnel are not doing patient imaging.
The present invention improves upon prior systems and methods, including, e.g., the systems and methods described in the following patents:                1. U.S. Pat. No. 7,071,474, entitled Methods and Apparatus for Tuning Scintillation Detectors;        2. U.S. Pat. No. 7,005,646, entitled Stabilized Scintillation Detector for Radiation Spectroscopy and Method; and        3. U.S. Pat. No. 6,835,935, entitled System and Method for Calibrating and Tuning a Gamma Camera.        
For reference, U.S. Pat. No. 6,835,935, assigned to the same assignee herein and incorporated herein by reference, teaches a system and method that is designed to calibrate and tune a gamma camera with minimal or no human intervention. The '935 system and method provides a valuable feature for the user in that the user is assured of optimal performance of the camera without requiring laborious procedures and expenditures of time that might otherwise be devoted to patient imaging. With the '935 system and method, analysis of PMT output pulses and calibration can be totally automatic. First, in one embodiment, the user does not need to handle a radioactive source because such sources can be made part of the camera. Second, the system computer can be programmed to monitor continuously the count rate and, thereby, determine when the system is being used and when the system is idle. When the system is idle (i.e., the count rate is approximately equal to the natural background plus the contribution of the radioactive sources), the system computer can automatically monitor and record individual PMT signals. When a sufficient number of data points have been stored for each PMT, the mean amplitude and variance of each tube's response to the events can be calculated. These calculated values may be compared to baseline values (e.g., values that were calculated at the time of the last tuning and calibration of the system, providing a database for comparison) and/or to calculated values associated with neighboring PMTs to determine whether and to what extent adjustments to the camera need to made. Further, the software may be programmed to analyze the results of the comparisons and automatically make the necessary PMT gain adjustments.
FIG. 1 is a diagram illustrating certain components of a gamma camera 100 according to one embodiment of the '935 system and method. As shown in FIG. 1, gamma camera 100 includes a scintillation crystal 102 (or “detector crystal 102”), a number of photomultiplier tubes (PMTs) 104(a) . . . (n), and a computer system 110 coupled to the output of each PMT 104. Advantageously, one or more very weak radioactive sources 106(a) . . . (n) is placed so as to be facing an entrance window side 103 of scintillation crystal 102 at fixed or known locations. Gamma camera 100 may also include a collimator 114 and a light guide 116. In one embodiment, sources 106 are positioned between collimator 114 and crystal 102. Data storage unit 112 stores data points for each PMT 104.
In one embodiment, sources 106 are positioned adjacent to the entrance window side 103 of scintillation crystal 102 at fixed or known locations. In a preferred embodiment, sources 106 are permanently or detachably affixed to entrance window side 103 of scintillation crystal 102 or to another component of camera 100, such as collimator 114. In a particular embodiment, a user of the camera 100 need not manually position sources 106 to occupy the fixed locations. For example, the sources may be pre-positioned and affixed to a component of camera 100 as part of the manufacturing process of the camera.
Sources 106 are chosen to have a photon energy that is below the source energies typical of diagnostic imaging, which are typically at least 140 keV. The source activity is also chosen to be below the limits set by regulatory agencies which would require licensing and inventory control. For example, Americium-241 (Am-241) emits a 60 keV X-ray and a long half-life. For activity levels less than 10 nCi, (nanocuries) such sources do not require radioactive material licenses.
Each radioactivity source 106, which is placed in a fixed location, causes scintillation photons to emanate from a small region directly “below” the source whenever an X-ray from the source 106 enters crystal 102. The scintillation photons produced by the X-rays will produce electronic signals of small amplitude in the photomultiplier tubes 104. Since the source activity is small, the probability of two absorption events overlapping in time is of negligible consequence.
The scintillation photons generated from each absorption event can be assumed to be located at a known point in the crystal 102 because each source 106 is placed in a fixed location and the range of the low energy photons (i.e., X-rays) within the scintillation crystal is short (e.g., <1 mm). Additionally, the mean number of scintillation photons produced from each X-ray absorption event will be near constant. Therefore, the signals produced in nearby PMTs, resultant from a number of scintillation photons generated from a single, mono energetic X-ray absorption and subsequently transported to the PMTs, will be random statistical variants about constant means, modified by any changes in light transport and PMT response and amplification (i.e. gain).
Because the PMT output signal caused by one of the sources 106 should be a random statistical variant about a constant mean, absent changes in light collection and absent changes in the PMT itself, a process 200, which is illustrated in FIG. 2 of the '935 patent, can be used to determine whether such changes have occurred and can be used tune PMTs 104 to compensate for the changes. Process 200 assumes a single source 106, but multiple sources may be used.
Although a number of methods of tuning gamma cameras are known, there continues to exist a need for improved systems and methods.